The invention relates generally to a surface coil element. It relates specifically to a quadrature surface coil adapted for use in a magnetic resonance imaging system.
It has been known that signal-to-noise ratio gains can be achieved by providing a quadrature surface coil which consists of a loop and a figure-eight, or butterfly, loop. The symmetry of the configuration of the quadrature surface coil results in the loops experiencing substantially zero mutual inductance, an important property in quadrature technology.
However, planar surface coils of this symmetrical quadrature configuration do not provide significant gains in signal-to-noise ratio.
It has been further known that the loop antenna used in coils is intrinsically a balanced structure, whereas the coaxial cable that transports the signal to the magnetic resonance imaging signal processor is adapted to function in an unbalanced manner, resulting in signal loss.
To minimize signal loss, it is desirable to make the proper balanced-to-unbalanced conversion between the coil and the coaxial cable. It has been known to perform this function with a circuit known as a balun. Baluns are typically constructed from loops of wire in either a transformer or split-inductor configuration.
However, baluns are known to be bulky and may require a finite core, which interferes with the highly magnetic environment of a magnetic resonance imaging system.
It has been further known to decouple a surface coil from the volume transmit coil, where the volume of the coil may comprise the head or body of the patient, while the volume coil is transmitting radio frequency pulses.
Decoupling is needed because many surface coils are designed to be in resonance at the precise frequency of transmission, which poses a serious health hazard to the patient, in that resonance effectively focuses the transmitted radio frequency energy onto a small area of patient tissue, causing a burn.
Further, decoupling is needed in that the surface coil will distort the transmitted radio frequency pulse, thereby causing inhomogeneities in the magnetic resonance image.
Decoupling of the surface coil from the volume coil is known to be best effected by including a decoupling circuit designed to open the loop during the radio frequency pulse.
Decoupling circuits are known to include high speed PIN diodes as the switching element because of their excellent radio frequency properties and their power handling capabilities. They are used in conjunction with inductors, capacitors, and/or quarter wavelength sections of transmission lines to open the circuit properly.
The known methods of switching PIN diodes on and off in decoupling circuits are active decoupling and passive decoupling.
Active decoupling uses a relatively small direct current bias provided by the magnetic resonance imaging signal processor to switch the diode, and is easy to implement. Passive decoupling uses the current induced in the loop by the radio frequency pulse to switch the diode, and is intrinsically safer in practice than active decoupling.
However, active decoupling results in a substantial risk to the patient, since many electrical components and interconnects are involved, any of which may fail.
Active decoupling uses PIN diodes having a large intrinsic region (I) between the P and N regions. The intrinsic region accumulates charge from the direct current bias, effectively allowing the small direct current bias to switch a much higher radio frequency current in the decoupling circuit.
However, while such a PIN diode works well in active decoupling, it is much too slow when used in passive decoupling. It requires a much higher turn-on voltage than is available in many surface coil applications, resulting in sub-optimal decoupling or no decoupling, with consequential potential harm to the patient.
A much faster PIN diode is required for passive decoupling, which calls for a diode having a much smaller intrinsic region than that used in active decoupling.
However, a smaller intrinsic region in the faster PIN diode results in significantly lower power handling capability.
Decoupling circuits are further known to have a large level of radio frequency current flowing through a tuned inductor-capacitor circuit, causing substantial radio frequency energy to radiate from the circuit.
However, since surface coils are positioned very close to the patient's tissue, the radiated energy is absorbed by the patient's tissue. This lowers the quality (Q) factor of the decoupler, thus lowering its performance, and can also result in undesirable image artifacts in the region that is absorbing the radiated radio frequency.
It has been further known that capacitive coupling occurs between the loop of the coil and the patient's tissue, since surface coils are placed so close to the patient. This results in an asymmetrical coil design that may produce undesirable inhomogeneities in the magnetic resonance image.
To minimize the effect of capacitive coupling, it has been known to use a Faraday shield around the loop to prevent or equalize the capacitive coupling to the patient, or multiple capacitors spaced evenly around the loop to minimize the effect of the capacitive coupling.
However, such known methods of minimizing the effect of capacitive coupling present technical difficulties, and result in a coil design that is more symmetrical during the receive phase of the imaging cycle, when the loop is closed, but not during the transmission phase, when the loop is opened by the decoupling circuit.
It is still further known in the art that there exists an optimal loop size and shape that maximizes the signal-to-noise ratio for a given region-of-interest that is a specific size, shape, and depth from the surface coil, and for a known precise source of noise in the magnetic resonance imaging system.
The optimal loop size and shape is such that making the size either smaller or larger and/or changing the shape or position will result in a lower signal-to-noise ratio.
For example, if the region of interest is very small and the source of noise is primarily thermal noise from the loop's radiation resistance, the optimal loop will be round in shape, while making the loop square will result in a slightly reduced signal-to-noise ratio. Further, if other sources of noise are also significant, the optimal loop size may vary to some degree, depending on the relative strengths of each noise source.
Many applications require images of regions that extend beyond the radius of the optimal loop size. A prime example is the thoracic region of the spine, in which case an oversized rectangular loop would be more suitable than a smaller round or square loop.
A larger loop having a lower signal-to-noise ratio, is known to be used for imaging a region that is larger than the optimal loop size. For example, for the case of the thoracic spine, it is known to use a rectangular loop that is about twice as long as it is wide, with a resulting signal-to-noise ratio loss of nearly six decibels compared to that of a square coil covering half the length.
An array of loops is known to be used to achieve the benefits of both maximal signal-to-noise ratio and sufficient coverage.
However, an array of loops, while technically feasible, requires multiple processors and elaborate digital processing to correctly reconstruct the image from the multiple loops, resulting in more expensive future systems, and not benefiting presently installed magnetic resonance imaging systems.